There have been broadly employed radiographic images such as X-ray images for diagnosis of the conditions of patients on the wards. Specifically, radiographic images using a intensifying-screen/film system have achieved enhancement of speed and image quality over its long history and are still used on the scene of medical treatment as an imaging system having high reliability and superior cost performance in combination. However, these image data are so-called analog image data, in which free image processing or instantaneous image transfer cannot be realized.
Recently, there appeared digital system radiographic image detection apparatuses, as typified by a computed radiography (also denoted simply as CR) and a flat panel detector (also denoted simply as FPD). In these apparatuses, digital radiographic images are obtained directly and can be displayed on an image display apparatus such as a cathode tube or liquid crystal panels, which renders it unnecessary to form images on photographic film. Accordingly, digital system radiographic image detection apparatuses have resulted in reduced necessities of image formation by a silver salt photographic system and leading to drastic improvement in convenience for diagnosis in hospitals or medical clinics.
The computed radiography (CR) as one of the digital technologies for radiographic imaging has been accepted mainly at medical sites. However, image sharpness is insufficient and spatial resolution is also insufficient, which have not yet reached the image quality level of the conventional screen/film system. Further, there appeared, as a digital X-ray imaging technology, an X-ray flat panel detector (FPD) using a thin film transistor (TFT), as described in, for example, the article “Amorphous Semiconductor Usher in Digital X-ray Imaging” described in Physics Today, November, 1997, page 24 and also in the article “Development of a High Resolution, Active Matrix, Flat-Panel Imager with Enhanced Fill Factor” described in SPIE, vol. 32, page 2 (1997).
To convert radiation to visible light is employed a scintillator panel made of an X-ray phosphor which is emissive for radiation. The use of a scintillator panel exhibiting enhanced emission efficiency is necessary for enhancement of the SN ratio in radiography at a relatively low dose. Generally, the emission efficiency of a scintillator panel depends of the phosphor layer thickness and X-ray absorbance of the phosphor. A thicker phosphor layer causes more scattering of emission within the phosphor layer, leading to deteriorated sharpness. Accordingly, necessary sharpness for desired image quality level necessarily determines the layer thickness.
Specifically, cesium iodide (CsI) exhibits a relatively high conversion rate of from X-rays to visible light. Further, a columnar crystal structure of the phosphor can readily be formed through vapor deposition and its light guide effect inhibits scattering of emitted light within the crystal, enabling an increase of the phosphor layer thickness.
However, the use of CsI alone results in reduced emission efficiency. For example, JP-B 54-35060 (hereinafter, the term JP-B refers to Japanese Patent Publication) disclosed a technique for use as an X-ray phosphor in which a mixture of CsI and sodium iodide (NaI) at any mixing ratio was deposited on a substrate to form sodium-activated cesium iodide (CsI:Na), which was further subjected to annealing as a post-treatment to achieve enhanced visible-conversion efficiency.
There were also proposed other means for enhancing light output, including, for example, a technique of rendering a substrate to form a scintillator thereon reflective, as described in patent document 1; a technique of forming a reflection layer on a substrate, as described in patent document 2; and a technique of a scintillator on a transparent organic film covering a reflective thin metal film provided on a substrate, as described in patent document 3. These techniques increased the light quantity but resulted in markedly deteriorated sharpness.
Techniques for placing a scintillator panel on the surface of a flat light-receiving element include, for example, those disclosed in JP-A Nos. 5-312961 and 6-331749. However, these are poor in production efficiency and cannot avoid deterioration in sharpness on a scintillator panel and a flat light-receiving surface.
Production of scintillator plates through a gas phase method were generally conducted by forming a scintillator layer on a rigid substrate such as aluminum or amorphous carbon and covering the entire surface of the scintillator layer with a protective layer, as described in, for example, patent document 4. However, formation of a scintillator layer on a substrate which cannot be freely bent is easily affected by deformation of the substrate or curvature at the time of vapor deposition when sticking a scintillator plate on the flat light-receiving element surface with paste, leading to defects such that uniform image quality characteristics cannot be achieved with the flat light-receiving surface of a flat panel detector. Further, a metal substrate exhibits a high X-ray absorption, producing problems, specifically when performing low exposure. In contrast, amorphous carbon which has been recently employed is useful in terms of low X-ray absorptivity but is difficult to be acceptable in practical production since no general-purpose one of a large size is available and its price is high. Accordingly, such problems have become serious along with the recent trend of increasingly larger flat panel detectors.
To avoid these problems was generally performed formation of a scintillator directly on the surface of a flat light-receiving element (i.e., on an imaging device) through vapor deposition or the use of a medical intensifying screen exhibiting flexibility but low sharpness instead of a scintillator plate. There was also disclosed the use of a soft protective layer such as poly-p-xylilene, as disclosed in, for example, patent document 5.
Although a scintillator material deposited directly on a flat light-receiving element exhibits enhanced image characteristics, such a scintillator material, however, has a cost defect such that a high-priced light-receiving element is wasted in occurrence of a defective product, and a heat treatment achieves enhancement of image characteristics of the scintillator material but a light-receiving element exhibits low heat resistance and is restricted in treatment temperature. There was also the problem of it being a cumbersome procedure such that it was necessary to include cooling the light-receiving element during thermal treatment.
Accordingly, there has been desired, to overcome the foregoing problems, development of a radiation flat panel detector which is superior in production suitability, inhibits deterioration of characteristics of a scintillator (or phosphor) layer during aging, protects the scintillator (or phosphor) layer from chemical deterioration or physical impact, results in little deteriorated sharpness in the scintillator plate and between flat light-receiving elements and results in uniform image quality characteristics.    Patent document 1: JP-B No. 7-21560    Patent document 2: JP-B No. 1-240887    Patent document 3: JP-A No. 2000-356679    Patent document 4: JP-B No. 3566926    Patent document 5: JP-A No. 2002-116258